Spectroscopy is a method for obtaining information on a molecular scale by the use of light. This information can be related to the rotational, vibrational and/or electronic states of the molecules probed as well as dissociation energy and more. The rotational and/or vibrational spectrum of a given molecule is specific for that molecule. As a consequence, molecular spectra in particular rotation and/or vibrational spectra are often referred to as ‘fingerprints’ related to a specific molecule. Information related to rotational, vibrational and/or electronic states of molecules can therefore be used to analyze a sample comprising a number of unknown molecular components, thereby obtaining knowledge about the molecular components in the sample.
The basis for a spectroscopic setup is a light source, e.g. a laser, which is used for illuminating a sample. The light from the light source (the incoming light) will interact with the sample, and often result in an alternation of the light which is transmitted through, emitted by, reflected by and/or scattered by the sample. By collecting the altered light and analyzing its spectral distribution, information about the interaction between the incoming light and the molecular sample can be obtained; hence information about the molecular components can be obtained.
The spectral distribution is typically measured by using a spectrometer. A spectrometer is an optical apparatus that works by separating the light beam directed into the optical apparatus into different frequency components and subsequently measuring the intensity of these components by using e.g. a CCD detector, a CCD array, photodiode or such.
The altered light reflecting interactions between the incoming light and the molecular sample can roughly be characterized as either emission or scattering. The emission signals have relatively broad spectral profiles as compared to scattering light signals, which normally display quite narrow spectral lines. One process often dominates over the other, but both processes can and most often will occur simultaneously. The intensity of the emitted light vs. the intensity of the scattered light depends among other things on the frequency and the power of the incoming light, the intensity of the incoming light at the measuring point in the sample, and the molecular components in the sample.
Scattered light can be classified as being either elastic or inelastic and these are characterized by being spectroscopically very narrow signals. Elastic scattering is referred to as Rayleigh scattering, in which there is no frequency shift. Rayleigh scattering thus has the same frequency as that of the incoming light.
The most commonly known example of inelastic scattering is Raman scattering, in which there is an energy interchanging between the molecule and the photons of the incoming light. The frequencies, i.e. the spectral distribution of the Raman scattered light will be different from that of the incoming light and uniquely reflect the specific vibrational levels of the molecule; hence it is a fingerprint spectrum. This can be used for identification of the molecular composition of the substance probed and/or the concentration of the specific molecules in the substance.
Raman scattering is a relatively weak process compared to e.g. Rayleigh scattering and fluorescence. Reduction of contributions from these other processes is thus desirable when collecting Raman scattered light. In addition, the intensity of the Raman scattered light depends strongly on the frequency and the intensity of the incoming light. If these are variable, it may therefore be essential to monitor power fluctuations in the incoming light if one is to receive reliable information about the distribution of molecular components in different samples and/or sample spot bases on analysis of the collected Raman scattered light, depending on the precision needed. The same is true if the analysis of the molecular components in a sample and/or different sample spots is bases on emission spectra.
Skin comprises a number of layers having different characteristics and containing different kinds of cells and structures. Various proposals for using Raman spectroscopy to measure glucose in skin or in other parts of the body have been made, but none of these has to date provided a system which can be used on most candidate subjects without adjustment to suit a particular individual and without calibration for that individual. It is thereby possible to calibrate an instrument against measurements of blood glucose concentration made on one individual or a group of individuals by other means such as chemical analysis and to apply that same calibration when the instrument is used on other individuals than the one or ones involved in the calibration. We have now appreciated that the key to achieving such a result is to ensure that the Raman scattered light that is collected for measurement originates at or close to a specific depth within the skin.
Caspers et al; Biophysical Journal, Vol 85, July 2003, describes an in vivo confocal Raman spectroscopy method and apparatus which is said to be useful for measuring glucose. It contains however no instruction as to the depth from which the Raman scattering should be collected in a glucose measurement and there is a strong suggestion deducible from the teaching that the apparatus had not actually been tried for this purpose.
WO2008/052221 describes a method and apparatus for coherent Raman spectroscopy that transmits light through a sample surface such as skin and tissue to a focal plane within the sample to measure for instance glucose. However, no teaching is present of the importance of selecting a particular depth for the focal plane or where this should be. Indeed, it is specifically acknowledged that using the described apparatus variations in the detected signal occur when the analyte concentration is constant due to effects of skin temperature and hydration. No suggestion is present that such effects can be avoided by a careful selection of the depth from which the measurements are taken.
WO97/36540 describes determination of the concentration of e.g. glucose using Raman spectroscopy and an artificial neural network discriminator. However, the Raman signals are not selectively obtained from a particular depth and the need to compensate for non-linearities arising from signals penetrating to a depth of >500 μm is discussed.
WO00/02479 discloses a method and apparatus for non-invasive glucose measurement by confocal Raman spectroscopy of the aqueous humor of the anterior chamber of the eye. Naturally, there is no teaching of a depth at which to make optimal measurements in skin.
WO2009/149266 refers back to Ermakov I V, Ermakova M R, McClane R W, Gellermann W. Opt Lett. 2001 Aug. 1; 26(15):1179-81, ‘Resonance Raman detection of carotenoid antioxidants in living human tissues.’ which describes using resonance Raman scattering as a novel noninvasive optical technology to measure carotenoid antioxidants in living human tissues of healthy volunteers. By use of blue-green laser excitation, clearly distinguishable carotenoid Raman spectra superimposed on a fluorescence background are said to be obtained.
Chaiken et al (Noninvasive blood analysis by tissue modulated NIR Raman spectroscopy, J. Chaiken et. al., Proc. of SPIE optical Eng., 2001, vol. 4368, p. 134-145) obtained a correlation of only 0.63 between Raman based measurements and fingerstick blood glucose measurements across several individuals, but were able to obtain a correlation of 0.90 for a single individual. The setup utilized by Chaiken et al comprises a collimated exitation beam and so naturally do not disclose any optimal focal depth.
WO2006/127766, WO02/07585 and US2006/0234386 all describe the use of Raman spectroscopy for measuring lactate through the skin surface. Lactate measurements may be used for various purposes including monitoring the effect of exercise and determining whether a person has died or is still living. In critical care, the monitoring of blood lactate is of importance. High levels of lactate may be associated with myocardial infarction, cardiac arrest, circulatory failure and emergency trauma situations.
The present invention now provides apparatus for non-invasive in vivo measurement by Raman spectroscopy of a substance, especially but not exclusively lactate or glucose but also including fatty acids, urea, carbamide, cholesterol and hemoglobin, present in interstitial fluid in the skin of a subject, comprising a light source, optical components defining a light path from said light source to a measurement location, a light detection unit, optical components defining a return path for Raman scattered light from said measurement location to said light detection unit, and a skin engaging member having a distal surface for defining the position of said optical components defining the return path with respect to a surface of said skin in use, and wherein said optical components defining a light path from said light source to a measurement location beneath the surface of the skin focus the light emitted from said source to a depth located at from 200 to 300 μm beneath the surface of the skin and optical components defining a return path for Raman scattered light selectively transmit to said light detection unit light scattered from near said measurement location such that at least 50% of Raman scattered light received at the light detection unit originates at depths from 60 to 400 μm beyond said distal surface of the skin engaging member.
The apparatus may include means for computing a concentration of a substance, particularly a metabolite, in interstitial fluid or blood based on analysis of said Raman scattered light. Metabolites may be glucose or lactate in particular, but also fatty acids, urea, carbamide, cholesterol, or hemoglobin. The Raman spectrum may be analysed by application thereto of a trained statistical model which relates peak intensities to the relevant metabolite concentration. This may be performed using partial least squares regression (PLS) as described in more detail in the references acknowledged in M. A. Arnold; In Vivo Near-Infrared Spectroscopy of Rat Skin Tissue with Varying Blood Glucose Levels; Anal. Chem. 2006, 78, 215-223 therein and in A. M. K. Enejder et al; Raman Spectroscopy for Non-invasive Glucose Measurements; Jnl of Biomedical Optics, 10(3), 031114; May/June 2005. Other forms of multivariate calibration may be used including Principal Component Analysis (PCA) in a manner analogous to that described in for instance A. G. Ryder, G. M. Connor and T. J. Glynn; Quantitative Analysis of Cocaine in Solid Mixtures using Raman Spectroscopy and Chemometric Methods; Journal of Raman Spectroscopy, 31; 221-227 (2000) or in J. T. Olesberg, L. Liu, V. V. Zee, and M. A. Arnold; In Vivo Near-Infrared Spectroscopy of Rat Skin Tissue with Varying Blood Glucose Levels; Anal. Chem. 2006, 78, 215-223. In general, statistical methods of spectrum analysis useful in calibrating detection of analytes from absorption spectra will be useful in analysis of Raman spectra also.
Preferably, said percentage is at least 55%. Preferably also, at least 90% of Raman scattered light received at the light detection unit originates at depths less than 600 μm beyond said distal surface of the skin engaging member. On the other hand, preferably less than 25% of Raman scattered light received at the light detection unit originates at depths less than 100 μm beyond said distal surface of the skin engaging member.
Preferably, at least 15% of Raman scattered light received at the light detection unit originates at depths from 200 to 300 μm beyond said distal surface of the skin engaging member.
Said optical components preferably defining a light path from said light source to a measurement location beneath a surface of skin preferably focus the light emitted from said light source to a depth located at from 210 to 300 μm, e.g. 250 beneath the surface of the skin.
In an alternative aspect, the invention provides apparatus of the kind described for measuring the concentration of a metabolite in interstitial fluid, wherein said Raman scattered light received at said detection unit includes at least light scattered by glucose or light scattered by lactate.
Apparatus according to the invention may comprise a hand piece for application to the skin containing components defining said measurement location in use, and one or more optical fibres connecting said hand piece to said light source and to a processing unit containing electronic circuitry for analysis of signals received from said light detection unit to provide said measurement therefrom.
The position distal of the skin engaging member of said measurement location is optionally adjustable and can be adjusted to be from 60 to 400 μm beyond said distal surface of the skin engaging member or can be adjusted to be from 200 (or 210) to 300 μm, beneath the surface of the skin. Alternatively, however the position distal of the skin engaging member of said measurement location is fixed, suitably such that the numerical parameters discussed above are achieved.
Thus, the depth of focus of the optical components defining said light path, and/or the optical components defining said return path may be fixed rather than adjustable.
The invention includes a method for non-invasive in vivo measurement by Raman spectroscopy of a said substance present in interstitial fluid in the skin of a subject, comprising directing light from a light source into the skin of said subject via optical components defining a light path from said light source to a measurement location in the skin, receiving Raman scattered light back from the skin at a light detection unit via optical components defining a return path for Raman scattered light from said measurement location to said light detection unit, whilst using a skin engaging member having a distal surface for defining the position of said optical components defining the return path with respect to a surface of said skin in use, and wherein said optical components defining a light path from said light source to a measurement location beneath a surface of skin focus the light emitted from said light source to a depth located at from 200 to 300 μm beneath the surface of the skin and said optical components defining a return path for Raman scattered light selectively transmit to said light detection unit light scattered from near said measurement location such that at least 50% of Raman scattered light received at the light detection unit originates at depths from 60 to 400 μm beyond said distal surface of the skin engaging member. The method is preferably performed using apparatus in accordance with the invention.
The method may include calibrating the output of the apparatus by the use of the apparatus to provide an output in respect of a known metabolite concentration prior to said measurement on said subject. Once calibrated the apparatus preferably is not calibrated again for a period of not less than a week, more preferably a month. Preferably, said calibration step of providing an output in respect of a known metabolite concentration is not carried out by the use of the apparatus on said subject.
Thus, the calibration may be conducted on a different subject for whom a blood glucose concentration or lactate or other metabolite concentration is known or may be conducted using a standard reference material such as a drop of metabolite solution placed in the measurement location or a solid phantom simulating a metabolite solution.
Any apparatus described herein may be used in such a method.
The invention further includes a handpiece for use in apparatus according to claim 1, said handpiece containing optical components defining a light path for light received at said handpiece from a light source to communicate said light to a measurement location, optical components defining a return path for Raman scattered light from said measurement location and for communicating said Raman scattered light to a remote light detection unit, and a skin engaging member having a distal surface for defining the position of said optical components defining the return path with respect to a surface of said skin in use, and wherein said optical components defining a light path from said light source to a measurement location beneath a surface of skin focus the light emitted from said light source to a depth located at from 200 to 300 μm beneath the surface of the skin and said optical components defining a return path for Raman scattered light selectively receive for communication to said light detection unit light scattered from near said measurement location such that at least 50% of Raman scattered light received at the light detection unit originates at depths from 60 to 400 μm beyond said distal surface of the skin engaging member.
The light source is preferably a laser. A preferred form of laser to use as the light source is a diode laser with a wavelength in the range of 300-1500 nm. Suitable preferred wavelengths are 785, 830, or 850 nm, 830 nm being especially preferred. A suitable power range is 50-1000 mW. For example, one may use a 830 nm, 500 mW FC-830 laser from RGB Lase.
The apparatus may include an optical probe for measuring light signals in which the optical components defining the light path from the light source to the measurement location comprise a first optical fiber guiding incoming light from said light source, a lens focusing said incoming light towards, i.e. into or onto, the measurement location. The optical components for defining a return path for Raman scattered light may comprise said lens and a distal portion of the said first optical fiber collecting altered light from the measurement location and a second optical fiber guiding the altered light to the light detection unit. However, instead of employing a second optical fiber as described, a spectrophotometer may be integrated directly into the handpiece. Optionally, there may be a further light detection unit (or light logging device) measuring intensity fluctuations in said incoming light, and this further light detection unit may advantageously be positioned after said first optical fiber, whereby said further light detection unit receives a part of said incoming light from said first fiber.
The use of optical fibers is advantageous in that although a microscope can be used, a microscope-based optical probe is not a movable object and a user's body part would be awkward to place in a position where measurements could be made. A possibility would be for the patient to insert his/her arm directly under or above the microscope objective in the microscope. Unfortunately, this is cumbersome if not impossible with most microscopes.
An optical probe employing not the whole microscope but only microscope objective(s) mounted separately on e.g. a table allows for a larger accessibility between probe and sample. Measurements of blood sugar, lactate or other metabolite levels in a patient in vivo become more convenient as the patients arm or finger can be placed in front of the microscope objective(s) without much difficulty. However, if the chosen sample is a leg, it might prove more difficult to place it appropriately in front of the microscope objective(s).
Inside the optical probe, said light logging device will normally be positioned after a dichroic mirror, which allows a minor part of the incoming light to either pass through the dichroic mirror and onto said light logging device or to be reflected by the dichroic mirror onto said light logging device. Alternatively, a splitting device can be positioned between said first fiber and said dichroic mirror, where said splitting device reflects a minor part of the incoming light onto said light logging device.
One advantage with using a light logging device is that it allows for a precise measure of the variations in the intensity of the incoming light at all material times. This ensures that variations in the intensity of the altered light due to variations in the incoming light and not sample variations can be compensated for.
In an embodiment of the invention, said lens focusing incoming light towards said sample is arranged at the surface of said optical probe such that said lens is in direct contact with the skin (213) during measuring.
An advantage with having the lens in direct contact with the skin during measurement is that the sample penetration depth, and thereby the distance from the optical probe to the sample focus point, is known exactly, as it is defined by the focal length of the lens.
In another embodiment of the invention, said optical probe further comprises a window, where said window is positioned between said lens and the skin, such that said window is in direct contact with the skin during measuring, and where the thickness of said window is smaller than the focal length of said lens.
An advantage with inserting a window between the lens and the skin is that it can provide an easier cleaning of the optical probe, if a fragile lens sensitive to cleaning is used.
Another advantage with inserting a window between the lens and the skin is that the penetration depth can be varied depending on the thickness of the window. This provides one way of setting the penetration depth to the value characterising the invention.
Equally, instead of having a solid window, a window aperture can be provided between the lens and the skin, the aperture being formed in the skin engaging member.
The optical probe according to the invention, may further comprise a dichroic mirror positioned after said first optical fiber, where said dichroic mirror reflects any percent between re_in=0 and 100 (e.g. 90%) and transmits any percent between tr_in=0 and 100 (e.g. 10%) of said incoming light, where re_in+tr_in=100 percent (ignoring losses), and reflects any percent between re_se=0 and 100 (e.g. 30%) and transmits any percent between tr_se=0 and 100 (e.g. 70%) of said altered light, where re_se+tr_se=100 percent (ignoring losses). Hence said dichroic mirror may reflect most of the incoming light and transmit most of the altered light.
Said dichroic mirror is normally positioned at an angle of 45 degrees in relation to the propagating direction of said incoming light out of said first optical fiber.
In an embodiment where most of the incoming light is reflected by the dichroic mirror, said light logging device may be positioned after said dichroic mirror, whereby said light logging device measures intensity fluctuations in said incoming light transmitted through said dichroic mirror.
In another embodiment where most of the incoming light is reflected by the dichroic mirror, a splitting device may be positioned between said first optical fiber and said dichroic mirror, whereby said light logging device measures intensity fluctuations in said incoming light reflected of by said splitting device.
In an embodiment of the invention, said dichroic mirror is transmitting most (e.g. ≧90%) of the incoming light whilst passing a minor portion (e.g. ≦10%) and is reflecting most of the altered light (e.g. ≧70%) whilst passing a smaller amount (e.g. ≦30%).
In an embodiment where most of the incoming light is transmitted by the dichroic mirror, said light logging device may be positioned after said dichroic mirror, whereby said light logging device measures intensity fluctuations in said incoming light reflected of by said dichroic mirror.
An advantage of having the light logging device situated directly after said dichroic mirror is that it utilizes the part of the incoming light, which is not reflected by the dichroic mirror, and otherwise would be lost. There is consequently no need for any additional optical components to be inserted inside the optical probe in order collect light for measuring of the fluctuations in the incoming light.
In one embodiment of the invention, the angle α between the direction (239) of light out of said first optical fiber (203) and the direction (241) of light entering said second optical fiber (227) is substantially α=90 degrees. The angle could also be in the range α=80-100 degrees.
In one embodiment of the invention, said optical probe further comprises at least a first aperture where said first aperture only allows altered light from the focus point in the skin to enter said second fiber thereby ensuring a confocal image, and where said first aperture is positioned immediately in front of said second fiber. Said aperture can be a separate element, but a narrow opening of said second fiber can equally well function as said aperture.
An advantage with using an optical aperture positioned before the second fiber is that the optical aperture works as a 3D depth filter eliminating optical signals generated outside of the confocal area, i.e. the sample focus spot. The advantage with using a confocal optical probe is that the altered light entering the second fiber arise solely from interactions between the incoming light and the skin at the focus spot; hence contributions from the cone-like areas above and below the focus spot are minimized or eliminated.
In another embodiment of the invention, one or more apertures can additionally be employed to obtain a sharper 3D depth image. A second aperture is preferably positioned between the skin and the lens focusing the light into the sample. This second aperture can be separate element, but a narrow opening of the optical probe at the point where light exits/is collected by the lens can equally well function as an aperture.
Although apparatus according to the invention is designed and configured for measuring optical signals in the skin in vivo, it could also be employed for measuring optical signals by immersing it into e.g. a blood sample thereby making the measurement in vitro.
Generally, the optical elements found inside an optical probe of apparatus according to the present invention are enclosed by a cover. A preferred optical probe can be moved around freely due to the use of flexible fibers for guiding light into and out of the optical probe. This enables easy in vivo measurements of e.g. blood sugar levels in a patient using different body areas such as an arm, a finger, a leg or similar. The apparatus may however be constructed so that the optical components are contained in a housing which defines a specific location on which to place a fingertip pad for performance of the measurement. The stratum corneum thickness of a fingertip pad will typically be from 10-40 μm (see Marks, James G; Miller, Jeffery (2006). Lookingbill and Marks' Principles of Dermatology (4th ed.). Elsevier Inc. Page 7. ISBN 1-4160-3185-5 and Thickness of the Stratum Corneum of the Volar Fingertips H. FRUHSTORFER, U. ABEL, C.-D. GARTHE, AND A. KNU” TTEL. Accordingly, the preferred measurement depths of 200-300 μm will be from 160 to 190 μm up to 260 to 290 μm below the stratum corneum. Depths of measurement for all skin areas are preferably from 50 to 390 μm, more preferably from 190 to 290 μm below the stratum corneum.
A primary application of the apparatus is generally to measure blood sugar or lactate levels in a patient. The level of glucose or lactate in blood correlates with the level in interstitial fluid at the selected depth.